Nanoengineered Osteoinductive and Elastomeric Scaffolds for Bone

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Nanoengineered Osteoinductive and Elastomeric Scaffolds for Bone Tissue Engineering Punyavee Kerativitayanan, Marco Tatullo, Margarita Khariton, Pooja Joshi, Barbara Perniconi, and Akhilesh K Gaharwar ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.7b00029 • Publication Date (Web): 03 Mar 2017 Downloaded from http://pubs.acs.org on March 7, 2017

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ACS Biomaterials Science & Engineering

Nanoengineered Osteoinductive and Elastomeric Scaffolds for Bone Tissue Engineering

Punyavee Kerativitayanana, Marco Tatullob,c, Margarita Kharitona, Pooja Joshia, Barbara Perniconib, and Akhilesh K. Gaharwara,d,e,* a

Department of Biomedical Engineering, Texas A&M University, College Station, TX 77841, USA b

c

d

Maxillofacial Unit, Calabrodental Clinic, 88900 Crotone, Italy

Regenerative Medicine Section, Tecnologica Research Institute, 88900 Crotone, Italy

Department of Materials Science and Engineering, Texas A&M University, College Station, TX 77841, USA

e

Center for Remote Health Technologies and Systems, Texas A&M University, College Station, TX 77843, USA *E-mail: [email protected]

Keywords: nanocomposites, poly(glycerol sebacate) (PGS), two-dimensional (2D) nanosilicates, bone regeneration, osteoinductive.

ACS Paragon Plus Environment

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ABSTRACT

This work reports synthesis and fabrication of porous and elastomeric nanocomposite scaffolds from

biodegradable

poly(glycerol

sebacate)(PGS)

and

osteoinductive

nanosilicates.

Nanosilicates are mineral-based two-dimensional (2D) nanomaterials with high surface area which reinforced PGS network. The addition of nanosilicates to PGS resulted in mechanically stiff and elastomeric nanocomposites. The degradation rate and mechanical stiffness of nanocomposite network could be modulated by addition of nanosilicates. Nanocomposite scaffolds supported cell adhesion, spreading and proliferation, and promoted osteogenic differentiation of preosteoblasts. The addition of nanosilicates to PGS scaffolds increased alkaline phosphatase (ALP) activity and production of matrix mineralization. In vivo studies demonstrated biocompatibility and biodegradable of nanocomposite scaffolds. Overall, the combination of elasticity and tailorable stiffness, tunable degradation profiles, and the osteoinductive capability of the scaffolds offer a promising approach for bone tissue engineering.


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INTRODUCTION Over 2.2 million bone grafting surgeries are performed worldwide each year, costing approximately 2.5 billion US dollars annually and is expected to rise due to aging population.1-2 Although, autograft is the clinical gold standard as it provides optimal osteogenesis, but results in donor site morbidity and post-operation complications including hernia formation, nerve injury, blood loss, and infection.3-5 Allograft is the surgeon’s second choice, but can results in immunogenic reactions and risk of disease transmission.6-10 Allograft processing, including freezing and irradiation, can mitigate these risks, but at the same time weaken its biomechanical and biochemical characteristics. Demineralized bone matrix (DBM) is another available option, but suffers from non-uniform processing procedures.6-10 Consequently, DBM is mostly used as a bone graft extender rather than a sole bone graft substitute. Due to above limitations, there is an unmet need to develop synthetic bone graft substitutes to promote bone regeneration. Macroporous bioactive ceramic granules are the most popular amongst synthetic bone grafts and bone graft substitutes. Surgeons mix these ceramic granules with patients’ blood and use as a putty to fill bone defects.6-10 Synthetic hydroxyapatite (sHAp), beta tricalcium phosphate (βTCP), and biphasic calcium phosphate (a mixture of sHAp and β-TCP) are most commonly used as ceramic granules.9 Compared to sHAp, bioactive glasses (Na2O-CaO-P2O5-SiO2) exhibit tenfold higher bioactivity index, indicating superior bone bonding ability. Nevertheless, translational success of these osteoconductive biomaterials has been restricted due to limited mechanical stability.10-12 Therefore, there is a demand for alternative materials such as nanocomposites to control and direct cell function for bone regeneration.13-16 In this regard, we recently reported that two-dimensional (2D) nanosilicates (Laponite Na+0.7[(Mg5.5Li0.3)Si8O20(OH)4]-0.7) induces osteogenic differentiation of human mesenchymal


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stem cells (hMSCs) without any osteoinductive supplements such as bone morphogenic protein2 (BMP-2).16-19 Osteoinductive properties of nanosilicates stem from their dissolution products, i.e., Na+, Mg2+, Li+, and Si(OH)4. Specifically, magnesium ions (Mg2+) were reported to promote adhesion of osteoblasts to material surfaces via fibronectin receptor α5β1 and β1 integrins, which brought about ossification and increased extracellular matrix protein expression.20-21 Lithium ions (Li+) activated the Wnt/β-catenin signaling pathway. This pathway activation resulted in increased cell proliferation and matrix mineralization. It also promoted osteogenic differentiation of mesenchymal cells while inhibiting osteoclastogenesis and apoptosis.22 Orthosilicic acid (Si(OH)4) stimulated osteogenic differentiation as well as promoted collagen type I synthesis.20, 23

Moreover, cytotoxic concentration of nanosilicates was ten-fold higher than that of sHAp,

suggesting that significantly higher dose of nanosilicates could be tolerated by cells.17 Possessing these properties, nanosilicates offer a growth-factor free approach for bone regeneration, minimizing complexities and expenses involved in growth factor delivery. Since nanosilicates have heterogeneous charge distribution, they interact strongly with hydrophilic polymers such as collagen and poly(ethylene glycol)(PEG)13-14 and have be developed as high performance elastomers24-27, moldable hydrogels28, injectable hemostats29, self-healing structures30, and drug delivery vehicles31-33. Our group recently reported that incorporation of nanosilicates in collage-based hydrogels gave rise to a four-fold increase in bone formation without any osteoinductive factors.19, 34 However, development of osteoinductive scaffolds with load-transducing mechanical properties remains a significant challenge. It is important to note that bone is a dynamic tissue where remodeling continuously takes place. Applied loads play critical roles in determining the rate of turnover as well as the formation of


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callus, its volume, and stiffness during bone healing. Thus, scaffolds with tailorable mechanical properties and degradation kinetics are of pivotal importance. Here,

we

report

the

fabrication

porous

poly(glycerol

sebacate)(PGS)/nanosilicates

nanocomposites scaffolds for bone tissue engineering. PGS is elastomeric and tough polyester, and undergoes surface erosion, making it appropriate for tissue engineering scaffolds.35-37 The degradation products are completely resorbable and have shown to trigger minimum inflammation and limited fibrous encapsulation compared to other synthetic polymers such as poly(lactic-co-glycolic

acid)(PLGA)

and

poly(lactic

acid)(PLA).35-37

PGS

is

also

osteoconductive and supports adhesion and differentiation of preosteoblasts in vitro.38 Since PGS has been shown to be osteoconductive, its action is expected to complement the osteoinductive capability of nanosilicates, resulting in enhanced bone regeneration. Furthermore, PGS could induce angiogenesis, contributing to its extensive studies as vascular graft.38-39 Since angiogenesis is generally coupled with osteogenesis, this property stresses its potential for bone tissue engineering.

MATERIALS AND METHODS Poly(glycerol sebacate)(PGS) Synthesis: Polycondensation of glycerol (Sigma-Aldrich) and sebacic acid (Sigma-Aldrich) in a 1:1 molar ratio was performed following published procedures to obtain PGS prepolymer.40-41 Briefly, equimolar ratios of glycerol and sebacic acid were allowed to react under nitrogen in a three-neck flask (round-bottom) at 120°C. The vacuum was turned (50 mTorr) on after 60 mins and the reaction was allowed to continue for 2 days. After the reaction was finished, the vacuum was turned off and polymer was cooled down to room temperature (under argon) and kept at 4°C for storage.


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Fabrication of PGS and PGS/Nanosilicates (PGS/nSi) Scaffolds: Porous scaffolds were fabricated by salt-leaching method. Briefly, NaCl (Sigma-Aldrich, ACS reagent grade) was grinded using a mortar and pestle, and then sieved to gather crystals that fall between 75 and 150µm diameter to use as porogens. VWR sieves #100 (150µm opening) and #200 (75µm opening) were used. Teflon molds were completely filled with sieved salts, and put into an oven overnight at 37°C. The Teflon molds were custom made to have a size of 6mm diameter and 2mm thickness. Next, in a glass vial, PGS was dissolved in 30%ethanol/70%chloroform (50%w/v), then 0%, 1%, 2.5%, 5%, and 10% of nanosilicates added. Nanosilicates (Laponite XLG) was obtained from BYK Additives Inc. Nanosilicates are 2D nanoclays with 20-50nm in diameter and 1nm in thickness with chemical formula – (Na+0.7[(Mg5.5Li0.3)Si8O20(OH)4]-0.7). A probe sonicator (FB120, Fisher Scientific) was used to mix nanosilicates in PGS solution until it turned homogeneous. The solution was added to the molds over the salt and allowed to penetrate between salt crystals until all the spaces were filled. The samples were kept in a fume hood for 48 hours or until the solvent completely evaporated, then another 24 hours in a vacuum desiccator for air bubble elimination. After that, the samples were thermally cured under vacuum at 130°C for 48 hours to covalently crosslink PGS prepolymer, similar to our previously published results.40-41 Then, the crosslinked samples were removed from the molds and submerged in deionized water under agitation to leach out the salt. The salt leaching process took 36 hours, with water changed every 4 hours for the first 12 hours. The scaffolds were lyophilized and stored in a vacuum desiccator until further studies. Surface and Cross-section Morphology: After successful fabrication, surface and cross-section images of the scaffolds were taken using a scanning electron microscope (SEM) (Neoscope JCM-5000). For cross-section imaging, the scaffolds were broken in half while submerged in


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liquid nitrogen in order to preserve their porous structure. The samples were kept in a vacuum desiccator for at least 24 hours for drying before being mounted on specimen stubs using a carbon tape. They were then sputter coated with gold/palladium before SEM imaging. Mechanical Properties: Mechanical properties were studied using cyclic compression testing (eXpert 7600, ADMET, USA) with the stain rate of 1mm/min until 60% strain for 8 cycles. From the stress-strain curves, compressive modulus, energy dissipation, and percent recovery upon reloading were calculated. Degradation Studies: Scaffolds were first weighed for initial dry weight (Wi). They were then submersed in 0.01M NaOH over a 1-week period at 37°C. After 1, 2, 4, 7 days, the samples were collected and lyophilized, and final dry weight (Wd) was measured. Equation 1 was used to calculate the loss in weight in percent given the initial weight (Wi) and final dry weight (Wd). Weight loss (%) =

ௐ೔ ିௐ೏ ௐ೔

‫ݔ‬100

(1)

Protein Adsorption: Fetal bovine serum (FBS) protein was used for protein adsorption study. Scaffolds were washed thoroughly with Dulbecco’s phosphate buffered saline (DPBS) before being incubated in 10% FBS at 37°C, 5% CO2, for 24 hours. Note that FBS purchased from Life Technologies was diluted with DPBS to make 10% FBS solution. After 24 hours, FBS was removed; and the scaffolds were washed trice with DPBS. The purpose of this washing step was to remove the proteins not being adsorbed on the scaffolds, so it should be gently done. After that, 2% sodium dodecyl sulfate (SDS) solution was added, and the well plate was shaken for 6 hours to detach adsorbed proteins. The solution was then collected for protein assay. Note that SDS solution (20%) purchased from Amresco was diluted by DPBS to make 2% SDS solution used in the study. Protein concentration was determined using a Micro BCATM Protein Assay Kit (Thermo Scientific) following manufacturer’s protocols.


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Cell Culture and In Vitro Cell Adhesion: MC-3T3 E1-subclone4 preosteoblasts (ATCC®) were cultured in alpha modification of Eagle’s medium (α-MEM, HycloneTM) supplemented with 10% FBS and 1% Pen/Strep antibiotic (Life Technologies), in humidified incubator at 37°C and 5% CO2 level. This medium would be referred to as normal growth media in this paper. The cells were harvested at 70% confluency for experiments. For sterilization, scaffolds were washed twice with DPBS then exposed to ultraviolet irradiation for 4 hours. The scaffolds were soaked in normal growth media in an incubator overnight prior to cell seeding to pretreat the surface which would subsequently facilitate cell attachment. It should be noted that for all cell studies, cell-seeded scaffolds were cultured in non-treated well plates to minimize cell adhesion to the well plate (i.e. maximize cell adhesion to the samples), whereas cells seeded on treated well plates (referred to as tissue culture plate or TCP here) were used as positive controls, and nonseeded TCP as negative controls. For cell adhesion study, the scaffolds were first glued to a glass slide using silicone sealant to minimize scaffold handling. Preosteoblasts were seeded at 1x105 cells in 5µl normal media per well on the scaffolds. After 3-hour incubation, additional 600µl of normal growth media was added. After 24 hours, normal media was replaced by osteoconductive media (α-MEM + 10mM β-glycerophosphate + 0.05mM ascorbic acid) in selected wells. Cell culture media was changed every 3 days. The samples were collected after 7 days in culture and prepared for SEM imaging. Samples were gently washed trice with DPBS before being fixed with 2.5% glutaraldehyde (Alfa Aesar®). Then, the samples were dehydrated using an ascending series of ethanol: 30%, 50%, 75%, 95%, and 100%, in order. Dehydrated samples were chemically dried using hexamethyldisilazane (HDMS) (electronic grade, Alfa Aesar®). The samples were left overnight


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in a fume hood to allow HDMS to completely evaporate. They were then mounted onto stubs, sputter coated, and imaged. In Vitro Cell Proliferation and Differentiation: Preosteoblasts were trypsinized at 70% confluency in culture and seeded on pre-treated scaffolds at approximately 5,000cells/100µ media. Cell proliferation was assessed using alamarBlue® assay (Thermo Scientific). The proliferation data was collected on days 1, 3, 5, 7, 10, 14 after the initial cell seeding. Osteogenic differentiation was assessed by investigating alkaline phosphatase (ALP) activity and matrix mineralization. Specifically, after 7 days in culture, the scaffolds were stained with BCIP/NBT substrate solution (Thermo Scientific) for ALP. The stained samples were imaged using a steromicroscope. Moreover, SensoLyte® pNPP Alkaline Phosphatase Assay Kit was used to quantify ALP activity. ALP activity was then normalized to dsDNA content for each sample. PicoGreen® dsDNA quantitation assay (PicoGreen® kit, Thermo Scientific) was performed using NanoDropTM 3300 fluorospectrometer. In addition, mineralization was assessed by Alizarin Red S (ARS) (2% solution, pH 4.2, Electron Microscopy Science) staining and quantification after 14 days in culture. Before staining, the samples were fixed with 2% glutaraldehyde, and then washed with excess DI water. ARS solution was added for 2 minutes. The samples were washed with excess DI water again before imaging. For quantification, 10% acetic acid (Fisher Scientific) was added to the stained samples under agitation for 30 minutes. The ARS-containing acetic acid was collected and neutralized with 10% ammonium hydroxide (Sigma-Aldrich) until the pH was 4.1-4.5. It was then read with a UV/VIS plate reader at OD405. Absorbance-ARS concentration conversion was done using a pre-plotted standard curve. In Vivo Study in Mice: Biocompatibility, biodegradability, and bioactivity of the scaffolds were studied in vivo. Eight-week-old non-immunodeficient male BALB/c mice were from Charles


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River. The animals were handled according to IACUC guidelines (Calabrodental Clinic, 88900 Crotone, Italy). Before operation, they were subjected to an intraperitoneal injection of Ketamin/xilezin (1ml/Kg of 10mg/ml mixed solution) anesthetic. PGS-10%nSi scaffolds were implanted in the left limb between a long bone and the associated muscle. Briefly, a pocket was created in the connective tissue between the femur and the adductor muscle. The scaffolds (2x4 mm) were inserted in the pocket, parallel to the femur bone, and transplanted in triplicate. The mice were hosted in standard conditions, and sacrificed after 3d, 6d, and 3w by cervical dislocation. Histology and histochemical analysis were performed on dissected grafts. The dissected grafts were embedded in tissue freezing medium (Leica, Wetzlar, Germany), and then frozen in cool isopropane. Lyica cryostat was used to obtain transverse cryosections of 7µm. The sections

were

then

strained

with

hematoxylin

and

eosin

(H&E),

toluidine

blue,

immunofluorescence, and alkaline phosphatase. Axioscop2 plus system equipped with Axicam HRc (Zeiss, Oberkochen, Germany) was used to take images. To quantify the fluorescence signal of the inflammatory cells within the scaffolds, we quantified the IF staining for CD206 by means of ImageJ software (https://imagej.nih.gov/ij/); data were expressed as integrated density, following the subtraction of the background autofluorescence. Statistics: The data was presented as mean ± standard deviation (n=3-5). Statistical analysis was performed by one-way analysis of variance (ANOVA) with Turkey’s post hoc test for pairwise comparison. Statistical significance designated with one, two, and three asterisks corresponded to p < 0.05, p < 0.01, and P < 0.001, respectively.


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RESULTS AND DISCUSSION Porous PGS and PGS-nanosilicates scaffolds were fabricated via salt-leaching method (Figure 1). Nanosilicates and PGS prepolymer were mixed in ethanol/chloroform (3:7) solvent using a probe sonicator. Then, the solution was added to salt-filled teflon molds. Salt (NaCl) was sieved to obtain particle sizes between 75µm and 150µm. Earlier studies have shown that pore size of ~ 100 µm are appropriate to facilitate cell migration and proliferation with a scaffold structure.42-43 Thermal curing at 130°C for 48hr yielded fully crosslinked PGS networks from PGS prepolymer. These crosslinked samples were then submerged in deionized water for salt leaching and to obtain porous scaffolds. Porous scaffolds were used to evaluate the structure, mechanical, and physiological characteristics. In vitro cell studies and in vivo experiments were also performed to evaluate their efficacy for tissue engineering applications.

Porous Microstructure of PGS/Nanosilicates Scaffolds The microstructure of salt-leached PGS and PGS-nanosilicate scaffolds were investigated using electron microscopy. The surface (lateral region) of PGS scaffold displayed porous networks (Figure 2a). The pore size ranged between 75-150µm as designed by the size of NaCl porogens. In addition, smaller micropores with the size of approximately 5-20µm were distributed between the macropores (75-150µm). It was likely that the micropores were generated during the curing process. Thermal curing of PGS involved trans-esterification reaction.44 This led to covalent crosslinking between polymer chains that resulted in 3D network of random coils, turning the viscous liquid into solid elastomer. The solvent and air bubbles were unlikely to be the sources of micropores as chloroform and ethanol are both highly volatile and solvent evaporation was performed under vacuum desiccation for at least 24 hours before curing


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to eliminate air bubbles. In addition, cross-section (transverse region) images showed that the pores were homogenously distributed across the width and depth of the scaffolds. Morphology of pores present on the scaffold surface and interior appeared to be similar, indicating a uniform microporous network. Microstructure of PGS and PGS-nanosilicate scaffolds also show porous network (Figure 2b). The pore size was similar amongst different scaffolds, suggesting that nanosilicate concentrations had no effects on pore size. This was expected since the pore size was mainly controlled by NaCl porogens. We could account this as an advantage of the salt-leaching method in comparison to other fabrication techniques; the resultant porous structure could be better controlled and was tunable independently of the scaffold composition. High magnification images showed that the pore walls of PGS scaffolds were relatively smooth, and the roughness increased upon the addition of nanosilicates. The increase in surface roughness could be attributed to the presence of nanosilicates and their interactions with PGS matrix. It was hypothesized that the rough topography would enhance cell adhesion and spreading.

Nanosilicates Increased Mechanical Stiffness without Compromising Elastomeric Properties Bone is a dynamic tissue and applied loads play vital roles in determining the rate of turnover as well as the formation of callus, its volume, and stiffness during bone healing. It was reported that scaffolds possessing some elasticity provided a load-transducing environment in which matrix deposition, new bone formation, and maturation could take place.38 PGS is elastomeric and reinforcing it with nanosilicates is expected to increase its strength and toughness. Mechanical characteristics of the scaffolds were investigated using uniaxial cyclic compression (8 cycles). All samples (both PGS and PGS/nanosilicates) exhibited non-linear stress-strain


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behavior (Figure 3a), which is typical for elastomeric materials. From the stress-strain curves, compressive modulus, recovery, and energy dissipation were determined. The results showed that the addition of nanosilicates significantly increases the compressive modulus of nanocomposites in a concentration-dependent manner (Figure 3b). The compressive modulus for the scaffolds containing 0%, 1%, 2.5%, 5%, and 10% nanosilicates were 27.9±5.1 kPa, 45.9±4.8 kPa, 67.2±16.1 kPa, 88.6±9.0 kPa, and 130.8±15.3 kPa, respectively. The addition of 10% nanosilicates to PGS enhanced the modulus for 4.5-fold compared to PGS scaffold. This could be because nanosilicates interacted strongly with PGS, restraining polymer chain movement as well as improving load transfer capability during deformation. In this regard, nanosilicates were reported to drastically improve mechanical properties of soft nanocomposites. 28, 45

We did not observe any significant change in modulus when samples were subjected to

multiple compressive cycles (1-8 cycles). To evaluate the elastomeric characteristics of nanocomposite we evaluated energy dissipation during each cycle was determined (Figure 3c). The first cycle displayed maximum energy absorption. It then decreased and remained rather unchanging during subsequent cycles for all scaffold compositions. For example, PGS-1%nSi scaffolds absorbed 1.8±0.3 kJ/m3 during cycle 1 and approximately 1.3±0.2 kJ/m3 during cycles 2-8. Moreover, the addition of nanosilicates does not compromise elasticity of scaffolds, as indicated by increased energy dissipation with increasing nanosilicate concentrations. During the first cycle, the scaffolds with 0%, 1%, 2.5%, 5%, and 10% nanosilicates absorbed 1.2±0.4 kJ/m3, 1.8±0.3 kJ/m3, 2.6±0.8 kJ/m3, 2.1±0.4 kJ/m3, and 3.1±1.6 kJ/m3, respectively. We observed that the addition of 10% nanosilicates to PGS network resulted in >2.5 times increase in energy dissipation. Our previous studies showed that the addition of 10% nanosilicates to solid PGS/nanosilicates nanocomposites gave rise to over


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6.5 times increase in energy dissipation.40 Since the porous structure of the scaffolds was controlled by porogens, it is likely that the effects of nanosilicates were less in comparison to their nonporous counterparts. Recovery of crosslinked network after cyclic mechanical deformation was calculated from the stress-strain curves. PGS network consists of random coils that are covalently linked with each other. The hydrogen bonding between –OH groups of the PGS backbone contribute to the observed elastomeric characteristics.35,

46

Interestingly, the addition of nanosilicates did not

compromise elasticity of the scaffolds despite that it enhanced strength and elasticity. In particular, nanosilicate concentration did not have statistically significant effects on network recovery (Figure 3c). For example, after 8 cycles of compression, recovery of PGS and PGS10%nSi scaffolds was found to be 98.4±1.4% and 97.6±3.4%, respectively. These results suggested that mechanical properties of PGS/nanosilicates scaffolds were tunable. The addition of nanosilicates increased mechanical strength, while maintaining their elastomeric properties. This would be beneficial as some flexibility is required in the early stage of fracture repair involving cartilage formation prior to bone calcification47 In the future, we aim to increase nanosilicate concentration or crosslinking density of PGS network to bring the mechanical properties closer to those of cancellous bone. Earlier studies have shown that nanocomposites containing 70% nanosilicates are cytocompatible.48 Other option to increase the mechanical stiffness is to decrease pore sizes scaffold or to incorporate crystalline polymer. Considering the achieved mechanical properties, it is promising that PGS/nanosilicates scaffolds could be used for craniofacial defects.

Nanosilicates Enhanced Physiological Stability of the Scaffolds


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It is well established that the degradation profile of scaffolds is profoundly important to how they perform. Scaffolds should give structural support and at the same time allow gradual replacement with newly formed bone. Under physiological conditions, PGS underwent surface erosion via cleavage of ester bonds.36 From a tissue engineering standpoint, surface-eroding polymers are more preferred than bulk-degrading ones since they can maintain more linear reduction in structural integrity and mechanical performances as mass loss continues over time. In preliminary studies, scaffold degradation was studied for one week in PBS, 37°C. Less than 3% weight loss was observed for all samples (data not shown). With standard deviation taken into account, the differences between compositions were statistically inconclusive. Under in vivo conditions, PGS degrade in approximately 35-50 days.36 In order to accelerate the degradation in few days, we carried out accelerated degradation test using 0.01M NaOH solution. Basic condition was used to accelerate the process (7-10 days), and it was anticipated that the trends would be more prominent.

Percentage of weight loss was measured on days 1, 2, 4, and 7. The addition of nanosilicates (1, 2.5, 5 and 10%) was found to significantly reduce weight loss compared to PGS scaffolds (Figure 4a). For example, after one week, PGS scaffolds were completely degraded, while nanocomposite scaffolds loaded with 1%, 2.5%, 5%, and 10% nanosilicates showed 47.7±24.4%, 60.7±14.2%, 34±6.6%, and 28±3.4% weight loss, respectively. That means more than 3.5 times reduction in weight loss upon the addition of 10% nanosilicates. Although, we observe a reduction in weight loss due to nanosilicate addition, but there was not significant difference between nanocomposites containing 1% and 2.5% nanosilicates. This might be attributed to significantly low amount of nanosilicates compared to PGS. The increases in physical integrity


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with the addition of nanosilicates could be observed for all studied time points. In addition, the scaffolds remained intact and maintained their shape upon degradation, i.e. they became smaller disks over time. This indicated a surface eroding nature, which was beneficial since the scaffolds would be able to better maintain their structural integrity during new tissue regeneration in comparison to bulk degrading scaffolds. It has been demonstrated that degradation kinetics of PGS could be entirely altered by varying degree of crosslinking. In this case, nanosilicates acted as crosslinkers, retarding degradation of polyester backbone resulting in enhanced physical integrity. Our previous studies showed that crosslinking density of thermally cured PGS/nanosilicates nanocomposites was increased upon the addition of nanosilicates.40 Sol-gel analysis was used to investigate degree of covalent crosslinking in that study. The sol (uncrosslinked) fraction was statistically significantly reduced with increasing nanosilicate concentrations. The results agreed with other published works that nanosilicates could serve as multifunctional crosslinkers.45,

49

It is anticipated that the

crosslinking involved transesterification with the secondary alcohol of glycerol. Nevertheless, the exact nature of crosslinking mechanisms will need further investigation. PGS degrades by hydrolysis of esters while nanosilicates slowly disintegrate at pH

Nanoengineered Osteoinductive and Elastomeric Scaffolds for Bone

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